Gammy ray and X-ray imaging detectors are used as part of imaging devices, e.g. positron emission tomography (PET) and single-photon emission computed tomography (SPECT) in medical and other applications. Such imaging detectors usually comprise a scintillation detector or a scintillator, e.g. a scintillator crystal or an array of scintillator crystals, coupled to a photodetector, e.g. an array of photosensitive elements. The scintillator scintillates, i.e. emits light flashes, in response to incoming, i.e. impinging, particles such as electrons, alpha particles, ions or high-energy photons. The emitted photons are captured by the photodetector, which, in turn, is read out by dedicated read-out electronics. Based on where and when scintillation photons are captured (i.e. the intensity distribution or spatial intensity distribution of the charges on the photodetector which may also be referred to as charge distribution), the temporal and spatial position and of the incident particles in the scintillator may be determined. Therefrom, an image representing this position may be generated. This image comprises information on where and when the respective particle was emitted, which can be exploited by medical or other imaging devices by introducing a substance emitting particles or by causing the emission of particles at certain areas in other ways. If, e.g., a patient is administered a radioactive tracer emitting a certain kind of particles (possibly in response to a metabolic reaction) an image can be generated as a representation of where these particles were emitted. Alternatively a gamma ray detector may also detect gamma rays emitted by a gamma ray source and interacting with an object (e.g. a patient) on their way to the detector.
One important issue in the context of such imaging approaches is the provided image resolution. This resolution depends on various factors such as the design of the photodetector or the scintillator (e.g. size or pitch of the photodetector array or the scintillator array in case arrays are used), the data processing and the used algorithms, the calibration of the different components, material properties (size, quality, . . . ), external conditions or other influences. The scintillator comprised in the detector may, e.g. comprise a single (monolithic) block, which results in a continuous distribution of the emitted scintillation photons (light distribution) in response to incident particles to be sampled and analyzed. Alternatively the scintillator may comprise an array of small crystal needles, which results in an intrinsic spatial resolution of the imaging detector given by the pitch of those needles. A higher number of crystal needles per area, i.e. smaller needles and/or more needles may, e.g., increase the resolution in that the positions of the incident particles can be determined with higher precision.
Thus, improving the intrinsic spatial resolution of such an imaging detector may be achieved by decreasing the pitch of the crystal elements in a scintillator array. Making the pitch of the crystal elements smaller, however, leads to a higher number of crystal elements that have to be correctly identified. There are two main strategies for the identification of the crystal being subject to the incident particle: either each individual scintillator crystal element can be individually read out by means of a dedicated photodetector element or the light-sharing method can be used. According to the light-sharing method, the pitch of the scintillator array is usually smaller than the pitch of the photo-detector array so that several crystal elements are placed over a single photodetector array pixel. In order to identify the respective crystal needles, i.e. the scintillator array element that was hit by the incident particle, it may then be evaluated how the scintillation light is distributed over multiple photodetector elements. In order to improve the detection and correct identification of the crystal needle a lightguide, i.e. an optical homogenous and transparent solid material, may be used for spreading the scintillation light over several photodetector array pixels. The distribution of the scintillation photons, i.e. the scintillation light or scintillation flash, over the photosensitive elements of the photodetector array may then be analyzed for identifying the scintillator array element that was hit. Further, the energy of the incident particle can be determined. However, extraction of the correct parameters (time, energy and position of the impact) is usually more difficult if light-sharing is used instead of individually reading out each scintillator array element. On the other hand, the required number of photosensitive elements in the photodetector (photodetector pixels) and the complexity of the data acquisition system may be reduced significantly, which may lead to lower device costs. For instance, modern clinical PET scanners have a number of scintillator crystal elements on the order of 104 to 105. The pitch of the arrays is usually approximately 4 mm leading to an intrinsic spatial resolution of about 4 mm. If each scintillator crystal array element is read out individually the same amount of photosensitive elements in the photodetector (photodetector pixels) and electronic channels would be required. The use of the light-sharing method can reduce the number of required photodetector pixels and electronic channels by an order or magnitude.
However, making use of the light-sharing method may also lead to disadvantages. For instance, in gamma ray detectors based on light-sharing the crystal, i.e. the crystal element, that is hit by an incident gamma ray and the energy of this gamma ray has to be extracted from a set of signals from all affected photosensitive elements in the photodetector, which usually requires an additional computation step. For positioning, the most widely used method is anger-positioning, i.e. the determination of the center-of-gravity or the centroid of the distribution. Anger-positioning is, however, heavily affected by missing signals, caused, e.g., by dead photosensitive elements in the photodetector or by the dead-time of one or more photosensitive elements. In Lerche et al., Maximum Likelihood Based Positioning and Energy Correction for Pixelated Solid State PET Detectors, Nuclear Science Symposium and Medical Imaging Conference Record, 2011, pp. 3027-3029, the authors present an alternative method for determining the position of an incident gamma ray and extracting the respective parameters. The approach is based on the Maximum Likelihood method. The most likely photo-conversion position in a scintillator array coupled to a photodetector array in light-sharing mode is determined by comparing the resulting light distribution with predetermined distributions for different photo-conversion positions in the scintillator. The most likely position, i.e. the position corresponding to the most similar light distribution, is used as an estimate for the photo-conversion position in the scintillator of the incident gamma ray. The authors show that the resolution of medical images may be improved by using the Maximum Likelihood position estimation method.
It is, however, not further detailed how the necessary reference distributions for the comparison are to be obtained.
In Yoshida et al., Calibration Procedure for a DOI Detector of High Resolution PET Through a Gaussian Mixture Model, IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 51, NO. 5, October 2004, a depth of interaction detector is developed for the next generation of positron emission tomography (PET) scanners. A statistical model based on the approach of a Gaussian mixture model (GMM) is introduced for crystal identification. The results of this method are used to generate a look-up-table.
In Ziock et al., 3D Millimeter Event Localization in Bulk Scintillator Crystals, IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 60, NO. 2, April 2013, a new technique to achieve a high level of performance through the use of close-coupled, coded-aperture shadow masks placed between the crystal and a position-sensitive phototransducer is presented.